This invention relates generally to nuclear magnetic resonance (NMR) imaging techniques. More particularly, it relates to methods for quantitatively measuring and compensation for eddy currents induced as a result of field gradient switching in magnetic resonance imaging (MRI) systems.
Magnetic Resonance Imaging (MRI), having its roots in magnetic resonance spectroscopy (MRS), has become a widely accepted and commercially available technique for obtaining digitized visual images representing the internal structure of objects (such as the human body) having substantial populations of atomic nuclei that are susceptible to nuclear magnetic resonance (NMR) phenomena. In MRI nuclei in a body to be imaged are polarized by imposing a strong main magnetic field Ho on the nuclei. Selected nuclei are excited by imposing a radio frequency (RF) signal at a particular NMR frequency. By spatially distributing the localized magnetic fields, and then suitably analyzing the resulting RF responses from the nuclei, a map or image of relative NMR responses as a function of the location of the nuclei can be determined. Following a Fourier analysis, data representing the NMR responses in space can be displayed on a CRT.
As shown in FIG. 1, an NMR imaging system typically includes a magnet 10 to impose the static magnetic field, gradient coils 14 for imposing spatially distributed magnetic fields along three orthogonal coordinates, and RF coils 15 and 16 to transmit and receive RF signals to and from the selected nuclei. The NMR signal received by the coil 16 is transmitted to a computer 19 which processes the data into an image displayed on display 24. The magnetic resonance image is composed of picture elements called xe2x80x9cpixels.xe2x80x9d The intensity of a pixel is proportional to the NMR signal intensity of the contents of a corresponding volume element or xe2x80x9cvoxelxe2x80x9d of the object being imaged. The computer 19 also controls the operation of RF coils 15 and 16 and gradient coils 14 through the RF amplifier/detector 21 and 22 and gradient amplifiers 20, respectively.
Only nuclei with odd number of protons and/or neutrons have a magnetic moment and thus are susceptible to NMR phenomena. In MRI, a strong static magnetic field is employed to align nuclei, generating a gross magnetization vector aligned in parallel to the main magnetic field at equilibrium. A second magnetic field, applied transverse to the first field as a single RF pulse, pumps energy into the nuclei, which causes the gross magnetization vector to flip by, for example, 90xc2x0. In certain instances this RF pulse is also followed by a second stronger RF pulse to cause the gross magnetization vector to flip by another 180xc2x0 (sometimes called a xe2x80x9cflopxe2x80x9d). After the excitation, the nuclei precess and gradually relax back into alignment with the static field. As the nuclei precess and relax, they will induce a weak but detectable electrical energy in the surrounding coils that is known as free induction decay (FID). These FID signals (and/or magnetic gradient-refocused field echoes thereof), collectively referred to herein as MR signals, are analyzed by a computer to produce images of the nuclei in space.
An operation whereby the various coils produces RF excitation pulses and gradient fields to result in and acquire an MR signal is called an acquisition pulse sequence. A graphical representation of a simple acquisition sequence is shown in FIG. 2. In this basic example, a gradient magnetic field, Tgrad, is first superimposed along the main field to sensitize a slice of nuclei to a particular RF frequency. Thereafter, a specific slice of nuclei are selected by a 90xc2x0 RF excitation pulse followed by a 180xc2x0 RF pulse to generate a MR signal which, loosely speaking, appears as an xe2x80x9cechoxe2x80x9d following the 180xc2x0 RF pulse. When the time between the 180xc2x0 RF pulse center and the detected MR signal center is the same as that between the 90xc2x0 RF pulse center and the 180xc2x0 RF pulse center, the MR signal produced is conventionally called a symmetric xe2x80x9cspin-echoxe2x80x9d signal and the particular timing of applied pulses and fields shown in FIG. 2 is known as a spin-echo sequence. In addition, the MR echo-signal center can be shifted by adjusting applied field gradients so that an asymmetric spin-echo (ASE) is produced instead.
The NMR frequency and the main B0 field are related by the Larmor relationship. This relationship, represented by the following equation, states that the angular frequency, xcfx890, of the precession of the nuclei is the product of the magnetic field, B0, and the so-called gyromagnetic ratio, xcex3, a fundamental physical constant for each nuclear species:
xcfx890=xcex3B0
In magnetic resonance imaging (MRI) and volume-localized magnetic resonance spectroscopy (MRS), switching of magnetic field gradients is an essential part of almost every pulse sequence. It is a well known fact that such action induces eddy currents (EC) in surrounding conductive materials in magnets and creates undesirable EC field gradients and B0 (main field) oscillations which ultimately cause image artifacts. Conventionally, some MRI and MRS systems employ a xe2x80x9cpre-emphasisxe2x80x9d network that is adjusted during a system alignment/calibration procedure in an attempt to minimize the effects of eddy-current induced gradient fields. Unfortunately, it is difficult to adjust the pre-emphasis network to effectively minimize the EC-induced gradient fields without having accurate measurements of the eddy-current induced fields.
Although, it is generally known that one can use small pick-up coils for directly measuring eddy currents, the use of pick-up coils for eddy current measurement typically requires additional hardware and special expertise in its use. For example, see xe2x80x9cA Method for Mapping Magnetic Fields Generated by Current Coilsxe2x80x9d, by J. Chankji, J. Lefevre and A. Briguet, J. Phys. E 18, 1014 (1985). Other somewhat less direct methods that rely on NMR measurement techniques to measure eddy currents-such as, for example, by measuring free induction decay (FID) signals of a small sample using small RF coils after switching off a field gradient pulsexe2x80x94are also known. See for example, xe2x80x9cGradient Time-Shape Measurement by NMRxe2x80x9d, by E. Yamamoto, H. Kohno, J. Phys. E 19, 708 (1986); xe2x80x9cA Simple Method of Measuring Gradient Induced Eddy Currents to Set Compensation Networksxe2x80x9d, by R. E. Wysong and I. J. Lowe, Magn. Reson. Med. 29, 119-121 (1993); xe2x80x9cQuantitative Characterization of the Eddy Current Fields in a 40-cm Bore Superconducting Magnetxe2x80x9d, by Q. Liu, D. G. Hughes, P. S. Allen, Magn. Reson. Med. 31, 73-76 (1994). Some NMR eddy current measurement methods even use a gross sample. See, for instance, xe2x80x9cTemporal and Spatial Analysis of Fields Generated by Eddy Currents . . . xe2x80x9d, by Ch. Boesch, R. Gruetter, and E. Martin, Magn. Reson. Med. 20, 268-284 (1991). Unfortunately, all of the above known methods are either very difficult to implement and/or lack the capability to distinguish between EC field gradients and B0 (main field) oscillationxe2x80x94both of which are associated with eddy currents but have significantly different effects on image quality.
It would be extremely useful for MRI systems to have the capability of conveniently measuring the eddy currents induced by field gradient switching without requiring special equipment, either for trouble-shooting system performance problems or for assisting set-up of a pre-emphasis arrangement for minimizing the detrimental effects of the eddy currents. Consequently, there is a need for a method and system for obtaining distinct quantitative measurements of both EC-induced field gradients and B0 oscillation in the presence of both large and small field inhomogeneities.
The present invention provides a method and apparatus for measuring and compensating the effects of eddy currents induced during NMR imaging operations. A cubic or cylindrical sample is placed in the imaging volume of a MRI system at a position centrally located with respect to the main magnetic field and oriented with its longitudinal axis parallel to a desired measuring direction. A magnetic field gradient pulse is applied for inducing eddy currents as well as for generation of a slice selective spin-echo signal. The spin-echo signal is acquired immediately after the termination of each eddy-current inducing gradient pulse. Two slices are selected along the desired measurement direction at equal symmetrical distances from the center of the main magnetic field. Two spin-echo signals are acquired for each slice with the eddy-current inducing gradient pulse reversed in polarity between the two echo signals. Quantitative values for eddy-current induced field gradients and B0 oscillations are then determined based on the precessing frequencies of the acquired NMR signals.
Image quality is improved by compensating for eddy currents effects by applying the quantified values of the field gradients and B0 oscillations to set an appropriate pre-emphasis network. In addition, gradient pulses in MRI/MRS acquisition sequences may be balanced to further account for gradient-induced eddy currents. Other aspects of the disclosed invention include measuring the time course of gradient switching, altering the signal acquisition sequences to measure eddy currents having long time constants, repeatedly measuring the eddy currents to assist in pre-emphasis adjustments, and measuring EC-induced field gradients and B0 oscillation in the presence of moderately large background field (main B0 field) inhomogeneities.